Mechanical Properties and In Vitro Biocompatibility of Hybrid Polymer-HA/BAG Ceramic Dental Materials

The aim of this study is to prepare hybrid polymer–ceramic dental materials for chairside computer-aided design/computer-aided manufacturing (CAD/CAM) applications. The hybrid polymer–ceramic materials were fabricated via infiltrating polymerizable monomer mixtures into sintered hydroxyapatite/bioactive glass (HA/BAG) ceramic blocks and thermo-curing. The microstructure was observed by scanning electron microscopy and an energy-dispersive spectrometer. The phase structure was analyzed by X-ray diffraction. The composition ratio was analyzed by a thermogravimetric analyzer. The hardness was measured by a Vickers hardness tester. The flexural strength, flexural modulus, and compressive strength were measured and calculated by a universal testing machine. The growth of human gingival fibroblasts was evaluated by a 3-(4,5-dimethylthiazol-2-yl)-2,5-diphenyl tetrazolium bromide (MTT) colorimetric assay and immunofluorescence staining. The results showed that the sintering temperature and BAG content affected the mechanical properties of the hybrid polymer–ceramic materials. The X-ray diffraction analysis showed that high-temperature sintering promoted the partial conversion of HA to β-tricalcium phosphate. The values of the hardness, flexural strength, flexural modulus, and compressive strength of all the hybrid polymer–ceramic materials were 0.89–3.51 GPa, 57.61–118.05 MPa, 20.26–39.77 GPa, and 60.36–390.46 MPa, respectively. The mechanical properties of the hybrid polymer–ceramic materials were similar to natural teeth. As a trade-off between flexural strength and hardness, hybrid polymer–ceramic material with 20 wt.% BAG sintered at 1000 °C was the best material. In vitro experiments confirmed the biocompatibility of the hybrid polymer–ceramic material. Therefore, the hybrid polymer–ceramic material is expected to become a new type of dental restoration material.


Introduction
Teeth defects, dentition defects, and missing teeth caused by caries, periodontal disease, and trauma have promoted the development of dentally restorative materials. In recent years, chairside computer-aided design/computer-aided manufacturing (CAD/CAM) has become an increasingly popular technology for personalized restorations. The machinable dental materials should possess good machinability, high strength, high cutting efficiency, and the ability to be easily polished [1].
Hydroxyapatite (HA) ceramics, biphasic calcium phosphate (BCP) ceramics, and bioactive glass (BAG) ceramics have been widely used in the biomedical field. BCP is composed of HA and β-tricalcium phosphate (β-TCP). BCP ceramics have proven to be adequately biocompatible, bioactive, and osteoconductive materials [2]. Compared with

Preparation of Hybrid Polymer-Ceramic Materials
Slurries were prepared by mixing deionized water with HA (5-20 µm), n wt.% (n = 10, 15, 20 and 25) of BAG (5 µm), 2.5 wt.% CMC-Na, and 2.5 wt.% PEG. The solid-liquid ratio was 1:3.5. The preparation of slurries referred to the articles of Tancred [18] and Chen [19]. Pure BAG and pure HA were also added with binders and dispersants as control groups, named 100 wt.% BAG and 0 wt.% BAG. The slurries were ball-milled with zirconia balls for 3.5 h at 200 r/min, passed through a 100-mesh sieve, and dried in a drying oven at 65 • C for 12 h. The dried powder was ground in a mortar. Molding processes included first compression molding (2 MPa, 30 s) and then cold isostatic pressing (250 MPa, 1 min). The green bodies were sintered at 1000 • C, 1100 • C, 1200 • C, and 1300 • C, with a heating rate of 2 • C/min and preservation time of 3 h. A total of 24 different ceramic blocks were prepared.
The polymerizable monomer mixtures were formulated with Bis-GMA and TEGDMA in a mass ratio of 1:1. A 1 wt.% BPO was used as the thermal initiator. They were stirred with a magnetic stirrer for 24 h until transforming into a homogeneous and low viscous liquid. The ceramic blocks were immersed in the monomer mixtures and placed in a vacuum-drying oven with a vacuum degree of −0.1 MPa for 3 days. Compared with normal atmospheric pressure, vacuum capillary action accelerated the complete infiltration of the monomer mixtures into the ceramic blocks. Finally, they were cured by hot isostatic pressing at 100 • C under a pressure of 150 MPa. Under the same conditions, the monomer mixtures were cured into pure resins as a control group. The process flow is shown in Figure 1.
with a heating rate of 2 °C/min and preservation time of 3 h. A total of 24 different cera blocks were prepared.
The polymerizable monomer mixtures were formulated with Bis-GMA TEGDMA in a mass ratio of 1:1. A 1 wt.% BPO was used as the thermal initiator. T were stirred with a magnetic stirrer for 24 h until transforming into a homogeneous low viscous liquid. The ceramic blocks were immersed in the monomer mixtures placed in a vacuum-drying oven with a vacuum degree of −0.1 MPa for 3 days. Comp with normal atmospheric pressure, vacuum capillary action accelerated the comp infiltration of the monomer mixtures into the ceramic blocks. Finally, they were cure hot isostatic pressing at 100 °C under a pressure of 150 MPa. Under the same conditi the monomer mixtures were cured into pure resins as a control group. The process f is shown in Figure 1. Cured hybrid polymer-ceramic blocks were cut to the appropriate dimensions cutting machine (SYJ-400, Shenyang Kejing Auto-instrument Co., Ltd., Shenyang, Ch equipped with an electroplated diamond saw blade. The spindle rotation speed of cutting machine was 1500 rpm, and the feed rate was 2 mm/min. The test specimens w polished with 500, 2000, and 4000 grit silicon carbide (SiC) abrasive papers in turn. Fin the dimensions of the cubic specimens were (25 ± 2) mm × (6 ± 0.1) mm × (3 ± 0.1) mm = 2), (25 ± 2) mm × (2 ± 0.1) mm × (2 ± 0.1) mm (n = 10) and (3 ± 0.1) mm × (3 ± 0.1) mm ± 0.1) mm (n = 10). After their cutting and polishing with 500 and 2000 grit SiC abra paper, cured pure resins, ceramic blocks, and cured hybrid polymer-ceramic blocks u in vitro experiments were (3 ± 0.1) mm × (3 ± 0.1) mm × (1 ± 0.1) mm (n = 12).

Microstructure and Phase Characterization
The cross-section of the specimens after cutting was treated with gold spraying. microstructure was observed by a field emission scanning electron microscope (FES Apreo 2S, Thermo Fisher Scientific, St. Louis, MO, USA). Elemental species and con analysis were differentiated by an energy dispersive spectrometer (Ultim MaX65, Oxf Wycombe, UK). Crystal structure analysis was performed using an X-ray diffractom (D8 Advance, Bruker, Karlsruhe, Germany).

Thermogravimetric Analysis
Proportion of polymers and ceramics in hybrid polymer-ceramic composites analyzed by a thermogravimetric analyzer (STA449F5, Netzsch, Selb, Germany) unde atmosphere at heating rate of 10 °C/min.

Microstructure and Phase Characterization
The cross-section of the specimens after cutting was treated with gold spraying. The microstructure was observed by a field emission scanning electron microscope (FESEM, Apreo 2S, Thermo Fisher Scientific, St. Louis, MO, USA). Elemental species and content analysis were differentiated by an energy dispersive spectrometer (Ultim MaX65, Oxford, Wycombe, UK). Crystal structure analysis was performed using an X-ray diffractometer (D8 Advance, Bruker, Karlsruhe, Germany).

Thermogravimetric Analysis
Proportion of polymers and ceramics in hybrid polymer-ceramic composites was analyzed by a thermogravimetric analyzer (STA449F5, Netzsch, Selb, Germany) under O 2 atmosphere at heating rate of 10 • C/min.

Porosity and Shrinkage
The porosity of the sintered ceramics was measured by the Archimedes drainage method. The shrinkage rate was calculated by using the volume of the ceramics before and after sintering. The ceramic specimens were not cut and polished. The dimensions of Polymers 2022, 14, 3774 4 of 20 specimens for both tests were (25 ± 2) mm × (10 ± 2) mm × (10 ± 2) mm (n = 5). Porosity was calculated by Equation (1): where m 1 is the mass of the dry ceramic blocks in the air; m 2 is the mass of saturated ceramic blocks in the water; m 3 is the mass of saturated ceramic blocks in the air.

Hardness
The Vickers hardness was tested by a Vickers hardness tester (VH1150, Wilson, MO, USA). The dimensions of specimens were (25 ± 2) mm × (6 ± 0.1) mm × (3 ± 0.1) mm (n = 2). A total of 10 points were randomly selected to apply the load on each cubic specimen surface. The mechanical properties of natural human teeth (Table 1) were used as a reference.

Flexural Strength and Flexural Modulus
The three-point bending test was performed on a universal testing machine (AGS, Shimadzu, Tokyo, Japan). According to ISO 10477: 2018 and ISO 4049: 2019 standards of dental restorative materials, 10 cubic specimens were prepared for each of the groups with the size of (25 ± 2) mm × (2 ± 0.1) mm × (2 ± 0.1) mm. The following formulas were used: for flexural strength (Equation (2)) for flexural modulus (Equation (3)) where F is the maximum load; (F/m) is the slope of the force-displacement curve; l is the distance between the supports (l = 20 mm); and b and h are the width and height, respectively, at the center of the specimen measured instantly prior to testing. The loading speed was 1 mm/min.

Compressive Strength
Tests were carried out on a universal testing machine (AGS, Shimadzu, Tokyo, Japan). 10 cubic specimens were prepared for each group with the size of (3 ± 0.1) mm × (3 ± 0.1) mm × (5 ± 0.1) mm. A (3 ± 0.1) mm × (3 ± 0.1) mm surface area was loaded with pressure. The compressive strength of the specimens was evaluated according to Equation (4): where F is the maximum load and S is the force area. The loading speed was 0.5mm/min.

In Vitro Cell Biocompatibility Measurements
To examine the cell viability of human gingival fibroblasts (HGFs, BT200003, Wuhan Biotower Biotechnology Co., Ltd., Wuhan, China) on the hybrid polymer-ceramic materials, direct contact experiment was performed by a 3-(4,5-dimethylthiazol-2-yl)-2,5- diphenyl tetrazolium bromide (MTT, HY-15924, MCE, Concord, CA, USA) colorimetric assay. HGFs were seeded in cell culture medium. The cell culture medium contained Dulbecco's modified Eagle medium (DMEM, PM150210, Procell Life Science & Technology Co., Ltd., Beijing, China), 10% fetal bovine serum (FBS, 111323, Shanghai Excell Biological Technology Co., Ltd., Shanghai, China), and 1% penicillin-streptomycin (PB180120, Procell Life Science & Technology Co., Ltd., China). HGFs were cultured in an incubator with 5% CO 2 at 37 • C. Cells were passaged using 0.25% trypsin (PB180225, Procell Life Science & Technology Co., Ltd., China). HGFs were seeded at 5 × 10 3 cells/well in 96-well cell culture plates containing the cubic specimens. The specimens (cured pure resins, 20 wt.% BAG ceramics sintered at 1000 • C, and hybrid polymer-ceramic materials of 20 wt.% BAG sintered at 1000 • C) were cleaned with ethanol and sterilized under ultraviolet light for 30 min. Their dimensions were (3 ± 0.1) mm × (3 ± 0.1) mm × (1 ± 0.1) mm (n = 12). The same culture medium without the specimen was used as a negative control. The growth status of cells was observed after 48 h and 72 h. After 72 h cell culture, 10 mL of MTT was added to each well and was incubated at 37 • C for 4 h. After that, the medium was removed. A total of 150 µL of DMSO was added and shaken for 10 min. The optical density (OD) was measured at a wavelength of 570 nm with a microplate reader (C22.2NO.1010.1, BioTek Instruments, Winooski, VT, USA). Results were expressed as a percentage of the negative control values.

Statistical Analysis
The results were statistically analyzed by SPSS software (SPSS 26.0, IBM Corp., Armonk, NY, USA) and compared using One-Way ANOVA and Tukey's honestly significant difference. Normality test using Shapiro-Wilk test. The significance level was 0.05. Figure 2a shows the green bodies after cold isostatic pressing. Figure 2b shows the sintered ceramics. Figure 2c shows the final specimens. Of all the observed specimen surfaces, the color changes of the pure HA (i.e., 0 wt.% BAG) groups were the most obvious. With the increase in the temperature, it became light gray, greyish-green, and gray, in turn. The color of the HA/BAG groups sintered at 1000 • C was light pink. The HA/BAG groups sintered at 1100 • C, the 10 wt.% and 15 wt.% BAG sintered at 1200 • C, and the 10 wt.% and 15 wt.% BAG sintered at 1300 • C were light yellow. The others were milky white. When the ceramics were sintered at 1000 • C and 1100 • C, uniform color distribution was observed over the cross-section of the corresponding composites. However, the periphery of the specimens at 1200 • C was light yellow and the interior was milky white. The specimens at 1300 • C were milky white on both the periphery and the interior.

Macromorphology and Microstructure
Comparing Figure 2a,b, the shrinkage of the specimens significantly increased with the increase in the sintering temperature. In Figure 2d, the crown restoration was carved from a hybrid polymer-ceramic block by a denture-processing machine (SD-5SH, Shanghai Shendiao, Shanghai, China). The milled edges were smooth, and no cracks or disintegration were observed. In Figure 2e, the test specimens of different sizes were cut from hybrid polymer-ceramic blocks and then polished. Figure 3 shows the microstructure of the fracture surfaces of the 20 wt.% BAG specimens with different sintering temperatures. The density of the porous ceramic was found to increase with the increasing of sintering temperature (Figure 3b,d,f,h). Finally, a dense ceramic was obtained at 1300 • C. Interconnected micropores and complete infiltration were observed in the specimens sintered at 1000 • C and 1100 • C (Figure 3a,c). Comparing Figure 3a,b and Figure 3c,d, the dark area represents the polymer phase and the bright area represents the ceramic phase. Figure 3i at a high magnification shows that the boundary between the polymer phase and the ceramic phase is blurred. A double network structure was formed by the close interconnection of the two phases. The specimen sintered at 1200 • C had more interconnected micropores in the periphery and fewer closed pores in the interior. The polymer phase was rarely observed in the interior. The specimens sintered at 1300 • C were relatively dense. The closed pores were smaller than the micropores of the specimens sintered at 1200 • C. The polymer phase was not observed whatsoever.
Polymers 2022, 14, x FOR PEER REVIEW 6 of 20 distribution was observed over the cross-section of the corresponding composites. However, the periphery of the specimens at 1200 °C was light yellow and the interior was milky white. The specimens at 1300 °C were milky white on both the periphery and the interior.
Comparing Figure 2a,b, the shrinkage of the specimens significantly increased with the increase in the sintering temperature. In Figure 2d, the crown restoration was carved from a hybrid polymer-ceramic block by a denture-processing machine (SD-5SH, Shanghai Shendiao, Shanghai, China). The milled edges were smooth, and no cracks or disintegration were observed. In Figure 2e, the test specimens of different sizes were cut from hybrid polymer-ceramic blocks and then polished.  Figure 3 shows the microstructure of the fracture surfaces of the 20 wt.% BAG specimens with different sintering temperatures. The density of the porous ceramic was found to increase with the increasing of sintering temperature (Figure 3b,d,f,h). Finally, a dense ceramic was obtained at 1300 °C. Interconnected micropores and complete infiltration were observed in the specimens sintered at 1000 °C and 1100 °C (Figure 3a,c). Comparing Figure 3a,b and Figure 3c,d, the dark area represents the polymer phase and the bright area represents the ceramic phase. Figure 3i at a high magnification shows that the boundary between the polymer phase and the ceramic phase is blurred. A double network structure was formed by the close interconnection of the two phases. The specimen sintered at 1200 °C had more interconnected micropores in the periphery and fewer closed pores in the interior. The polymer phase was rarely observed in the interior. The specimens sintered at 1300 °C were relatively dense. The closed pores were smaller than the micropores of the specimens sintered at 1200 °C. The polymer phase was not observed whatsoever. Figure 2. Macromorphologies of the specimens: (a) green bodies after cold isostatic pressing, (b) ceramic blocks after high-temperature sintering, (c) specimens after monomer mixtures' infiltration and curing, (d) hybrid polymer-ceramic material dental crown with 20 wt.% BAG sintered at 1000 • C, and (e) cubic test specimens after cutting. Figure 4 shows the energy spectrum of the hybrid polymer-ceramic material with 20 wt.% BAG sintered at 1000 • C. Si represents the BAG and C represents the polymer. Ca represents the ceramic phase, which is contained in both the HA and BAG. It can be seen from Figure 4 that the distribution of the HA, BAG, and polymer is relatively uniform. In addition, the energy spectrum shows that the C content of 20 wt.% BAG groups sintered at 1000-1300 • C are 32.85 wt.%, 28.09 wt.%, 19.08 wt.%, and 15.75 wt.%. No polymer phases were found in the interior of the ceramics sintered at 1200 • C and in the ceramics sintered at 1300 • C. Therefore, the specimens at 1200 • C and 1300 • C are not hybrid polymer-ceramic materials.

X-ray Diffraction (XRD) Analysis
When the sintering temperature increased, the diffraction peak of HA gradually weakened, the content of HA decreased, and the content of tricalcium phosphate (TCP) increased (Figure 5c-e). There are three polymorphic forms of TCP. The β phase appears below 1180 • C. The α phase appears at 1180-1400 • C. If the α phase is cooled slowly, it will change to the β phase. The α' phase appears above 1470 • C [24,25]. Comparing the three allotropes, β-TCP has good biocompatibility and mechanical strength. Its degradation rate is lower than that of α-TCP [26]. The XRD patterns of the ceramics were compared with the β-TCP standard card. The characteristic peaks and the sub-strength peaks of the two were consistent. Characteristic peaks of α-TCP phase were not found, indicating that only β-TCP was generated under this experimental condition. In addition, a new phase, Ca2SiO4, was also generated.  from Figure 4 that the distribution of the HA, BAG, and polymer is relatively uniform. In addition, the energy spectrum shows that the C content of 20 wt.% BAG groups sintered at 1000-1300 °C are 32.85 wt.%, 28.09 wt.%, 19.08 wt.%, and 15.75 wt.%. No polymer phases were found in the interior of the ceramics sintered at 1200°C and in the ceramics sintered at 1300 °C. Therefore, the specimens at 1200 °C and 1300 °C are not hybrid polymer-ceramic materials.

X-ray Diffraction (XRD) Analysis
When the sintering temperature increased, the diffraction peak of HA gradually weakened, the content of HA decreased, and the content of tricalcium phosphate (TCP) increased (Figure 5c-e). There are three polymorphic forms of TCP. The β phase appears below 1180 °C. The α phase appears at 1180-1400 °C. If the α phase is cooled slowly, it will change to the β phase. The α' phase appears above 1470 °C [24,25]. Comparing the three allotropes, β-TCP has good biocompatibility and mechanical strength. Its degradation rate is lower than that of α-TCP [26]. The XRD patterns of the ceramics were compared with the β-TCP standard card. The characteristic peaks and the sub-strength peaks of the two were consistent. Characteristic peaks of α-TCP phase were not found, indicating that only β-TCP was generated under this experimental condition. In addition, a new phase, Ca2SiO4, was also generated. Figure 5a,b are ceramics with the same sintering temperature (1000 °C). When the BAG content increased from 10 wt.% to 20 wt.%, the degree of conversion from HA to β-TCP decreased. This is similar to the experimental result of Tancred [18].  Figure 6 displays the thermogravimetric (TG) curves of the HA/BAG ceramic hybrid polymer-ceramic materials with different sintering temperatures and diffe BAG contents. The quality of the HA/BAG ceramic with 10 wt.% BAG sintered at 100 hardly changed with the temperature. Compared with the ceramic, the hybrid polym ceramic materials' curves had obvious steps and a substantial decline in quality. The cu polymers in the hybrid polymer-ceramic materials underwent a thermal cracking reac with the increase in the temperature. In the hybrid polymer-ceramic material wit  Figure 5a,b are ceramics with the same sintering temperature (1000 • C). When the BAG content increased from 10 wt.% to 20 wt.%, the degree of conversion from HA to β-TCP decreased. This is similar to the experimental result of Tancred [18].  Figure 6 displays the thermogravimetric (TG) curves of the HA/BAG ceramic and hybrid polymer-ceramic materials with different sintering temperatures and different BAG contents. The quality of the HA/BAG ceramic with 10 wt.% BAG sintered at 1000 • C hardly changed with the temperature. Compared with the ceramic, the hybrid polymerceramic materials' curves had obvious steps and a substantial decline in quality. The cured polymers in the hybrid polymer-ceramic materials underwent a thermal cracking reaction with the increase in the temperature. In the hybrid polymer-ceramic material with 10 wt.% BAG sintered at 1000 • C, bound water accounted for 1.21%, ceramic for 63.18%, and polymer for 35.61%. In the hybrid polymer-ceramic material with 20 wt.% BAG sintered at 1000 • C, ceramic accounted for 74.97% and polymer for 17.9%. In the hybrid polymer-ceramic material with 20 wt.% BAG sintered at 1100 • C, ceramic accounted for 83.96% and polymer for 12.38%. This illustrated that with the increase in the BAG content or sintering temperatures, the proportion of ceramics increases and the proportion of polymers decreases in the hybrid polymer-ceramic materials.  Figure 6 displays the thermogravimetric (TG) curves of the HA/BAG ceramic hybrid polymer-ceramic materials with different sintering temperatures and diffe BAG contents. The quality of the HA/BAG ceramic with 10 wt.% BAG sintered at 100 hardly changed with the temperature. Compared with the ceramic, the hybrid polym ceramic materials' curves had obvious steps and a substantial decline in quality. The cu polymers in the hybrid polymer-ceramic materials underwent a thermal cracking reac with the increase in the temperature. In the hybrid polymer-ceramic material wit wt.% BAG sintered at 1000 °C, bound water accounted for 1.21%, ceramic for 63.18%, polymer for 35.61%. In the hybrid polymer-ceramic material with 20 wt.% BAG sint at 1000 °C, ceramic accounted for 74.97% and polymer for 17.9%. In the hybrid polym ceramic material with 20 wt.% BAG sintered at 1100 °C, ceramic accounted for 83.96% polymer for 12.38%. This illustrated that with the increase in the BAG content or sinte temperatures, the proportion of ceramics increases and the proportion of polym decreases in the hybrid polymer-ceramic materials.

Surface Roughness
The dimensions of the three materials after polishing with 500 and 2000 grit SiC abrasive papers were (3 ± 0.1) mm × (3 ± 0.1) mm × (1 ± 0.1) mm (n = 12). The materials were cured pure resin, ceramic with 20 wt.% BAG sintered at 1000 • C, and hybrid polymerceramic material with 20 wt.% BAG sintered at 1000 • C (Figure 7). Typical grooves and damage are visible on the surface of the materials. The grain direction is in the polishing direction. It can be seen from Table 2 that the order of the surface roughness of the different materials prepared by the same cutting machine and the same abrasive papers is as follows: cured pure resin > hybrid polymer-ceramic material > ceramic. were cured pure resin, ceramic with 20 wt.% BAG sintered at 1000 °C, and hybrid polymer-ceramic material with 20 wt.% BAG sintered at 1000 °C (Figure 7). Typical grooves and damage are visible on the surface of the materials. The grain direction is in the polishing direction. It can be seen from Table 2 that the order of the surface roughness of the different materials prepared by the same cutting machine and the same abrasive papers is as follows: cured pure resin > hybrid polymer-ceramic material > ceramic.   Figures 8 and 9 show the shrinkage and porosity of the ceramics with different sintering temperatures and different BAG contents. With the increase in the sintering temperature or BAG content, the shrinkage gradually increased, and the porosity gradually decreased. The 0 wt.%, 15 wt.%, 20 wt.%, and 25 wt.% BAG ceramics sintered at 1200 °C, and the 0-25 wt.% BAG ceramics sintered at 1300 °C, had a high shrinkage and extremely low porosity, which means that the monomer mixtures could not penetrate the ceramics. Combined with the above studies, the specimens sintered at 1000 °C and 1100 °C, 10 wt.%, and 100 wt.% BAG sintered at 1200 °C were hybrid polymer-ceramic materials. Therefore, the shrinkage of the HA/BAG ceramic areas of the hybrid polymerceramic materials was 5.12-33.71%, and the porosity was 15.84-42.82%. The HA/BAG ceramic with the highest porosity was the 15 wt.% BAG sintered at 1000 °C, and the lowest was the 10 wt.% BAG sintered at 1200 °C.  Porosity   Figures 8 and 9 show the shrinkage and porosity of the ceramics with different sintering temperatures and different BAG contents. With the increase in the sintering temperature or BAG content, the shrinkage gradually increased, and the porosity gradually decreased. The 0 wt.%, 15 wt.%, 20 wt.%, and 25 wt.% BAG ceramics sintered at 1200 • C, and the 0-25 wt.% BAG ceramics sintered at 1300 • C, had a high shrinkage and extremely low porosity, which means that the monomer mixtures could not penetrate the ceramics. Combined with the above studies, the specimens sintered at 1000 • C and 1100 • C, 10 wt.%, and 100 wt.% BAG sintered at 1200 • C were hybrid polymer-ceramic materials. Therefore, the shrinkage of the HA/BAG ceramic areas of the hybrid polymer-ceramic materials was 5.12-33.71%, and the porosity was 15.84-42.82%. The HA/BAG ceramic with the highest porosity was the 15 wt.% BAG sintered at 1000 • C, and the lowest was the 10 wt.% BAG sintered at 1200 • C.

Shrinkage and
Polymers 2022, 14, x FOR PEER REVIEW 11 Figure 8. Shrinkage of ceramics with different sintering temperatures and different BAG conte the same lowercase letters indicate that there was no significant difference between different contents at the same sintering temperature (p = 0.05). A the same uppercase letters indicate that t was no significant difference between different sintering temperatures with the same BAG con (p = 0.05). The data were significantly drawn from a normally distributed population at the level.  Shrinkage of ceramics with different sintering temperatures and different BAG conte the same lowercase letters indicate that there was no significant difference between different contents at the same sintering temperature (p = 0.05). A the same uppercase letters indicate that t was no significant difference between different sintering temperatures with the same BAG con (p = 0.05). The data were significantly drawn from a normally distributed population at the level.

Figure 9.
Porosity of ceramics with different sintering temperatures and different BAG conten the same lowercase letters indicate that there was no significant difference between different contents at the same sintering temperature (p = 0.05). A the same uppercase letters indicate that t was no significant difference between different sintering temperatures with the same BAG con (p = 0.05). The data were significantly drawn from a normally distributed population at the level.

Hardness
The hardness of the hybrid polymer-ceramic materials generally increased with increase in the sintering temperature or BAG content ( Figure 10). The uninfiltr HA/BAG ceramics sintered at 1200 °C and 1300 °C exhibited a higher flexural stren and flexural modulus compared to the other groups related to the high density of ceramics. The hardness of the hybrid polymer-ceramic material for 10-25 wt.% HA/B groups ranged from 0.89 to 1.96 GPa, and the hardest was found for the 25 wt.% BA 1100 °C. In the control groups, the hardness of the 100 wt.% BAG sintered at 1100 °C the greatest (3.51 ± 0.09 GPa).

Figure 9.
Porosity of ceramics with different sintering temperatures and different BAG contents. a the same lowercase letters indicate that there was no significant difference between different BAG contents at the same sintering temperature (p = 0.05). A the same uppercase letters indicate that there was no significant difference between different sintering temperatures with the same BAG content (p = 0.05). The data were significantly drawn from a normally distributed population at the 0.05 level.

Hardness
The hardness of the hybrid polymer-ceramic materials generally increased with the increase in the sintering temperature or BAG content ( Figure 10). The uninfiltrated HA/BAG ceramics sintered at 1200 • C and 1300 • C exhibited a higher flexural strength and flexural modulus compared to the other groups related to the high density of the ceramics. The hardness of the hybrid polymer-ceramic material for 10-25 wt.% HA/BAG groups ranged from 0.89 to 1.96 GPa, and the hardest was found for the 25 wt.% BAG at 1100 • C. In the control groups, the hardness of the 100 wt.% BAG sintered at 1100 • C was the greatest (3.51 ± 0.09 GPa).

Flexural Strength and Flexural Modulus
The flexural strength of the hybrid polymer-ceramic materials for 10-25 wt.% BAG groups ranged from 57.61 to 93.38 MPa (Figure 11). In the control groups, the hybrid polymer-ceramic material with 100 wt.% BAG sintered at 1100 • C had the largest flexural strength (118.05 ± 9.84 MPa). With the increase in the BAG content, the flexural strength of the hybrid polymer-ceramic materials increased at first and then decreased, reaching a maximum when the BAG content was 20 wt.%. With the increase in the sintering temperature (1000 • C→1100 • C), the flexural strength of the hybrid polymer-ceramic materials decreased. As shown in Figure 12, the flexural moduli of all the hybrid polymerceramic materials ranged from 20.26 to 39.77 GPa. The hybrid polymer-ceramic material with 25 wt.% BAG sintered at 1100 • C had the largest flexural modulus. The flexural strength and flexural modulus of the uninfiltrated HA/BAG ceramics at 1200 • C and 1300 • C was high. This was related to the high density of the ceramics.

Compressive Strength
The posterior tooth crown has an extremely complex geometry and bears a great chewing force. The compressive strength test is similar to the chewing behavior. As shown in Figure 13, the compressive strengths of all the hybrid polymer-ceramic materials ranged from 60.36 to 390.46 MPa. The hybrid polymer-ceramic material with 25 wt.% BAG sintered at 1100 • C had the greatest compressive strength. With the increase in the BAG content, the shrinkage rate of the ceramics increased and the compressive strength of the hybrid polymer-ceramic materials also increased.
Polymers 2022, 14, x FOR PEER REVIEW 12 Figure 10. Hardness of specimens with different sintering temperatures and different BAG con . a the same lowercase letters indicate that there was no significant difference between different contents at the same sintering temperature (p = 0.05). A the same uppercase letters indicate that t was no significant difference between different sintering temperatures with the same BAG con (p = 0.05). The data were significantly drawn from a normally distributed population at the level.

Flexural Strength and Flexural Modulus
The flexural strength of the hybrid polymer-ceramic materials for 10-25 wt.% B groups ranged from 57.61 to 93.38 MPa (Figure 11). In the control groups, the hy polymer-ceramic material with 100 wt.% BAG sintered at 1100 °C had the largest flex strength (118.05 ± 9.84 MPa). With the increase in the BAG content, the flexural stren of the hybrid polymer-ceramic materials increased at first and then decreased, reachi maximum when the BAG content was 20 wt.%. With the increase in the sinte temperature (1000 °C→1100 °C), the flexural strength of the hybrid polymer-cera materials decreased. As shown in Figure 12, the flexural moduli of all the hybrid polym ceramic materials ranged from 20.26 to 39.77 GPa. The hybrid polymer-ceramic mat with 25 wt.% BAG sintered at 1100 °C had the largest flexural modulus. The flex strength and flexural modulus of the uninfiltrated HA/BAG ceramics at 1200 °C and °C was high. This was related to the high density of the ceramics. Figure 10. Hardness of specimens with different sintering temperatures and different BAG content. a the same lowercase letters indicate that there was no significant difference between different BAG contents at the same sintering temperature (p = 0.05). A the same uppercase letters indicate that there was no significant difference between different sintering temperatures with the same BAG content (p = 0.05). The data were significantly drawn from a normally distributed population at the 0.05 level. Figure 10. Hardness of specimens with different sintering temperatures and different BAG co . a the same lowercase letters indicate that there was no significant difference between different contents at the same sintering temperature (p = 0.05). A the same uppercase letters indicate that was no significant difference between different sintering temperatures with the same BAG co (p = 0.05). The data were significantly drawn from a normally distributed population at the level.

Flexural Strength and Flexural Modulus
The flexural strength of the hybrid polymer-ceramic materials for 10-25 wt.% groups ranged from 57.61 to 93.38 MPa (Figure 11). In the control groups, the hy polymer-ceramic material with 100 wt.% BAG sintered at 1100 °C had the largest flex strength (118.05 ± 9.84 MPa). With the increase in the BAG content, the flexural stre of the hybrid polymer-ceramic materials increased at first and then decreased, reach maximum when the BAG content was 20 wt.%. With the increase in the sinte temperature (1000 °C→1100 °C), the flexural strength of the hybrid polymer-cer materials decreased. As shown in Figure 12, the flexural moduli of all the hybrid polym ceramic materials ranged from 20.26 to 39.77 GPa. The hybrid polymer-ceramic mat with 25 wt.% BAG sintered at 1100 °C had the largest flexural modulus. The flex strength and flexural modulus of the uninfiltrated HA/BAG ceramics at 1200 °C and °C was high. This was related to the high density of the ceramics. Figure 11. Flexural strength of specimens with different sintering temperatures and different BAG contents. a the same lowercase letters indicate that there was no significant difference between different BAG contents at the same sintering temperature (p = 0.05). A the same uppercase letters indicate that there was no significant difference between different sintering temperatures with the same BAG content (p = 0.05). The data were significantly drawn from a normally distributed population at the 0.05 level. Figure 11. Flexural strength of specimens with different sintering temperatures and different contents. a the same lowercase letters indicate that there was no significant difference betw different BAG contents at the same sintering temperature (p = 0.05). A the same uppercase le indicate that there was no significant difference between different sintering temperatures wit same BAG content (p = 0.05). The data were significantly drawn from a normally distrib population at the 0.05 level. Figure 12. Flexural modulus of specimens with different sintering temperatures and different contents. a the same lowercase letters indicate that there was no significant difference betw different BAG contents at the same sintering temperature (p = 0.05). A the same uppercase le indicate that there was no significant difference between different sintering temperatures wit same BAG content (p = 0.05). The data were significantly drawn from a normally distrib population at the 0.05 level.

Compressive Strength
The posterior tooth crown has an extremely complex geometry and bears a g chewing force. The compressive strength test is similar to the chewing behavior. As sh in Figure 13, the compressive strengths of all the hybrid polymer-ceramic mate ranged from 60.36 to 390.46 MPa. The hybrid polymer-ceramic material with 25 w BAG sintered at 1100 °C had the greatest compressive strength. With the increase in BAG content, the shrinkage rate of the ceramics increased and the compressive stre of the hybrid polymer-ceramic materials also increased. Figure 12. Flexural modulus of specimens with different sintering temperatures and different BAG contents. a the same lowercase letters indicate that there was no significant difference between different BAG contents at the same sintering temperature (p = 0.05). A the same uppercase letters indicate that there was no significant difference between different sintering temperatures with the same BAG content (p = 0.05). The data were significantly drawn from a normally distributed population at the 0.05 level. Figure 12. Flexural modulus of specimens with different sintering temperatures and different contents. a the same lowercase letters indicate that there was no significant difference betw different BAG contents at the same sintering temperature (p = 0.05). A the same uppercase le indicate that there was no significant difference between different sintering temperatures wit same BAG content (p = 0.05). The data were significantly drawn from a normally distrib population at the 0.05 level.

Compressive Strength
The posterior tooth crown has an extremely complex geometry and bears a g chewing force. The compressive strength test is similar to the chewing behavior. As sh in Figure 13, the compressive strengths of all the hybrid polymer-ceramic mate ranged from 60.36 to 390.46 MPa. The hybrid polymer-ceramic material with 25 w BAG sintered at 1100 °C had the greatest compressive strength. With the increase in BAG content, the shrinkage rate of the ceramics increased and the compressive stre of the hybrid polymer-ceramic materials also increased. Figure 13. Compressive strength of specimens with different sintering temperatures and different BAG contents. a the same lowercase letters indicate that there was no significant difference between different BAG contents at the same sintering temperature (p = 0.05). A the same uppercase letters indicate that there was no significant difference between different sintering temperatures with the same BAG content (p = 0.05). The data were significantly drawn from a normally distributed population at the 0.05 level.

Biocompatibility
The human gingival fibroblasts (HGFs) adhered and proliferated well on the cured pure resin, HA/BAG ceramics, and hybrid polymer-ceramic materials after the 48 h culture.
They had a typical elongated fusiform morphology. Over time, the number and coverage of the HGFs after the 72 h culture increased, and the arrangement was compact (Figure 14). No obvious round, floating cells and dead cells were observed. Short-term exposure of HGFs to the polymer did not change the cell morphology. The cell viability of the HGFs after the 72 h culture was measured by the MTT assay (Figure 15), and a statistical difference in the cell proliferation was found between the surfaces of the different materials. The cell viability of the hybrid polymer-ceramic material group was 90.78% of that of control group, which was between the polymer group (94.09%) and ceramic group (81.56%). Figure 16 shows the immunofluorescence staining of HGFs on different materials. The size and morphologies of the nuclei in the four groups were similar. The cytoskeleton of the HGFs in the control group and polymer group was vivid and spread out. In the ceramic group and the hybrid polymer-ceramic material group, the shape of the HGFs was narrow and the cytoskeleton staining was not clear. The growth morphology of HGFs in the ceramic group was the worst.
different BAG contents at the same sintering temperature (p = 0.05). the same uppercase letters indicate that there was no significant difference between different sintering temperatures with the same BAG content (p = 0.05). The data were significantly drawn from a normally distributed population at the 0.05 level.

Biocompatibility
The human gingival fibroblasts (HGFs) adhered and proliferated well on the cured pure resin, HA/BAG ceramics, and hybrid polymer-ceramic materials after the 48 h culture. They had a typical elongated fusiform morphology. Over time, the number and coverage of the HGFs after the 72 h culture increased, and the arrangement was compact ( Figure 14). No obvious round, floating cells and dead cells were observed. Short-term exposure of HGFs to the polymer did not change the cell morphology. The cell viability of the HGFs after the 72 h culture was measured by the MTT assay (Figure 15), and a statistical difference in the cell proliferation was found between the surfaces of the different materials. The cell viability of the hybrid polymer-ceramic material group was 90.78% of that of control group, which was between the polymer group (94.09%) and ceramic group (81.56%). Figure 16 shows the immunofluorescence staining of HGFs on different materials. The size and morphologies of the nuclei in the four groups were similar. The cytoskeleton of the HGFs in the control group and polymer group was vivid and spread out. In the ceramic group and the hybrid polymer-ceramic material group, the shape of the HGFs was narrow and the cytoskeleton staining was not clear. The growth morphology of HGFs in the ceramic group was the worst.

Biocompatibility
The human gingival fibroblasts (HGFs) adhered and proliferated well on the cu pure resin, HA/BAG ceramics, and hybrid polymer-ceramic materials after the culture. They had a typical elongated fusiform morphology. Over time, the number coverage of the HGFs after the 72 h culture increased, and the arrangement was com ( Figure 14). No obvious round, floating cells and dead cells were observed. Short-t exposure of HGFs to the polymer did not change the cell morphology. The cell viab of the HGFs after the 72 h culture was measured by the MTT assay ( Figure 15), a statistical difference in the cell proliferation was found between the surfaces of different materials. The cell viability of the hybrid polymer-ceramic material group 90.78% of that of control group, which was between the polymer group (94.09%) ceramic group (81.56%). Figure 16 shows the immunofluorescence staining of HGF different materials. The size and morphologies of the nuclei in the four groups w similar. The cytoskeleton of the HGFs in the control group and polymer group was v and spread out. In the ceramic group and the hybrid polymer-ceramic material gr the shape of the HGFs was narrow and the cytoskeleton staining was not clear. growth morphology of HGFs in the ceramic group was the worst.

Discussion
Enamel and dentin are considered orthotropic due to their microstructural feat [27]. In order to achieve a "dental-like" effect, dental restorative materials can recapitu the biomimetic morphology of tooth tissue [28]. From the micro view (Figures 3 an the ceramic phase and polymer phase of the experimental material are distinct. Fro macro view, the ceramic particles and polymers are randomly dispersed, so the mat is homogeneous. As in almost all materials used for dental restoration, the hybrid poly ceramic material is isotropic. The sintering temperature significantly affects the mechanical properties of hy polymer-ceramic materials. An the experiment showed, with the increase in temperature (1150 °C→1350 °C), the microhardness and flexural strength of the HA/ ceramics increased first and then decreased [29]. For hybrid polymer-ceramic mate the ratio between the ceramic and polymer also influences the mechanical prope With the increase in the sintering temperature, the shrinkage of the HA/BAG cera increases, the porosity decreases, the penetration of monomer mixtures decreases, and proportion of ceramics in the composites increases. The HA/BAG ceramics sintere 1000 °C and 1100 °C have interconnected pores that allow for the easy infiltration o monomer mixtures, which can be used to prepare hybrid polymer-ceramic mate However, the ceramics sintered at 1200 °C and 1300 °C had glass phases and were de This is similar to the experimental results of Chen [19]. The locally closed micropor dense structure make the monomer mixtures impermeable. So, they are not hy polymer-ceramic materials.
The BAG content also influences the mechanical properties of hybrid polym ceramic materials. SiO2 and Si are key components for the surface activity and biolo properties of BAG. They play important roles in the decomposition of HA [30]. experiment of Li [31] showed that adding a small amount of SiO2 (0.5→1→5 w

Discussion
Enamel and dentin are considered orthotropic due to their microstructural features [27]. In order to achieve a "dental-like" effect, dental restorative materials can recapitulate the biomimetic morphology of tooth tissue [28]. From the micro view (Figures 3 and 4), the ceramic phase and polymer phase of the experimental material are distinct. From a macro view, the ceramic particles and polymers are randomly dispersed, so the material is homogeneous. As in almost all materials used for dental restoration, the hybrid polymer ceramic material is isotropic.
The sintering temperature significantly affects the mechanical properties of hybrid polymer-ceramic materials. An the experiment showed, with the increase in the temperature (1150 • C→1350 • C), the microhardness and flexural strength of the HA/BAG ceramics increased first and then decreased [29]. For hybrid polymer-ceramic materials, the ratio between the ceramic and polymer also influences the mechanical properties. With the increase in the sintering temperature, the shrinkage of the HA/BAG ceramics increases, the porosity decreases, the penetration of monomer mixtures decreases, and the proportion of ceramics in the composites increases. The HA/BAG ceramics sintered at 1000 • C and 1100 • C have interconnected pores that allow for the easy infiltration of the monomer mixtures, which can be used to prepare hybrid polymer-ceramic materials. However, the ceramics sintered at 1200 • C and 1300 • C had glass phases and were dense. This is similar to the experimental results of Chen [19]. The locally closed micropores or dense structure make the monomer mixtures impermeable. So, they are not hybrid polymer-ceramic materials.
The BAG content also influences the mechanical properties of hybrid polymer-ceramic materials. SiO 2 and Si are key components for the surface activity and biological properties of BAG. They play important roles in the decomposition of HA [30]. The experiment of Li [31] showed that adding a small amount of SiO2 (0.5→1→5 wt.%) promoted the significant conversion of HA to TCP, while adding 5 wt.% and 10 wt.% SiO 2 had little difference with respect to the amount of HA converted into TCP. Padilla's experiment confirmed that in 20 wt.% BAG ceramic sintered at 1300 • C, HA was significantly reduced and β-TCP was the main phase [32]. These are similar to the results of our experiment. The hardness, flexural strength, and compressive strength of the hybrid polymer-ceramic material of the pure HA (i.e., 0 wt.% BAG) were all low. With the addition of BAG, the mechanical properties were all improved. This was attributed to the decrease in the porosity and the formation of amorphous glass [33].
For commercial machinable dental materials, the hardness of zirconia ceramics such as Lava TMY-TZP (3M ESPE) is about 13 GPa [34]; the hardness of alumina ceramics (VITA In-Ceram Alumina, Vita Zahnfabrik, Bangkok, Thailand) is about 19.8 GPa; and the hardness of polymethyl methacrylate Esters (Polident PMMA, Polident, Volčja Draga, Slovenia) is about 0.25 GPa [35]. If the material with too high of a hardness and Young's modulus, such as metal, is used, it will lead to the abrasion of the opposite and adjacent teeth [36,37]. In turn, the lack of contact between the two teeth interferes with the efficiency of chewing [38]. If the hardness of the material is too low, the abrasion resistance and scratch resistance will be poor. Thus, such a material can only be used as a temporary restorative material. The hardness of hybrid polymer-ceramic materials in this experiment is relatively appropriate, which is between enamel (2.7-6.4 GPa) and dentin (0.12-0.67 GPa) ( Table 1).
Flexural strength is meaningful for repairing the masticatory function of defective and missing teeth. Especially for fixed bridge restorations or patients with abnormal occlusal habits, the flexural strength of the materials should be as high as possible [39]. The flexural strength of the hybrid polymer-ceramic materials prepared in this experiment matches that of enamel (60-90 MPa) (  [35]. Their flexural modulus is close to the elastic modulus of dentin . Their compressive strength is close to that of dentin (230-370 MPa) ( Table 1). High modulus values are a typical characteristic of stiff materials, such as metal and glass. However, polymers and porous structural materials are generally at low values of this parameter [13]. The flexural modulus of hybrid polymer-ceramic materials sintered at 1100 • C was higher than that sintered at 1000 • C, since high-hardness ceramics were more dominant in the hybrid polymer-ceramic composites. The flexural modulus of the HA/BAG groups at 1300 • C was very large, because the dense ceramics were not easy to deform.
The processing, finishing, and polishing of restorations affects not only the mechanical properties and aesthetics (surface staining and reflectance), but also the health of the soft tissue (plaque accumulation and gingival irritation), and the marginal integrity of the restoration [40]. In this experiment, the same processing and polishing methods were used for different materials, and the surface roughness of the materials was different ( Table 2). Using a different finishing technique on the same material also results in inconsistencies in roughness. A rougher surface will increase the adhesion of bacteria, such as S. sanguinis [41], one of the bacteria that colonizes the teeth early on, and Candida albicans [42], which causes oral infections. Besides roughness, other surface properties (surface charge and surface energy) also influence biofilm formation [43]. There is no further exploration of the material surface properties in this experiment. Therefore, it is necessary to find a more suitable processing and finishing technique for hybrid polymer-ceramic materials.
At present, commercial resin-based CAD/CAM materials contain monomers such as bisphenol-A-glycidyl dimethacrylate (Bis-GMA), triethylene glycol dimethacrylate (TEGDMA), urethane dimethacrylate (UDMA), or bisphenol-A-ethoxylated-glycidyl dimethacrylate (Bis-EMA). In our experiment, Bis-GMA and TEGDMA were selected as organic phases. Bis-GMA has the stable chemical structure of an ether bond, which can prevent the hydrolysis of saliva esterase [44]. The small molecule diluent TEGDMA reduces the viscosity of Bis-GMA. However, Dental materials inevitably interact and cause adverse biological impacts. These include various chemicals in saliva (such as solvolysis, hydrolysis, and alcoholysis) and physical reactions (such as abrasion and erosion) [45]. After exposure to Bis-GMA with a concentration for 50% of the maximal effect (EC50) for 6 h, human gingival fibroblasts (HGFs) were induced to DNA double-strand breaks [46]. TEGDMA also induces mitochondrial damage and oxidative stress in HGFs [47]. Another experiment showed that TEGDMA did not induce DNA strand breaks of HGFs at concentrations up to 10 mM until 24 h [48]. The reason is that the eluted monomers are unlikely to reach the concentrations that may induce necrotic cell death under physiological conditions [49]. According to the data provided by the manufacturers and experiments, the monomers released from commercial resin composite products were below the estimated daily intake limits [50][51][52].
Human gingiva consists of overlying epithelial structures and the underlying connective tissue. Human gingival fibroblasts (HGFs) are the most abundant cells in connective tissue and are very sensitive to the surrounding environment [53]. In addition, HGFs are easy to extract and have strong proliferation. Their culture success rate and resistance were higher than that of human gingival keratinocytes (HGKs) [54]. Studies showed hybrid polymer-ceramic materials were affected HGFs more than HGKs [55,56]. There are also 3D-printed dental restoration and CAD/CAM materials that are used to evaluate biocompatibility through the adhesion and proliferation of HGFs [57,58].
In order to detect the cytotoxicity of our hybrid polymer-ceramic material, the growth of HGFs was tested by the MTT method and immunofluorescence staining. We expected that the HGFs' morphology was typically spindle-shaped with many long cell processes that explore the environment and communicate with other cells [59]. The results showed that the fusiform HGFs grown on a hybrid polymer-ceramic material presented a wellstretched morphology and a proliferative state. The cell growth of the ceramic group was poor, which was similar to the experimental result of Ruano [60]. In addition to the differences in material types, several studies showed that the surface roughness of the materials influences cell attachment and growth [61]. The surface roughness of the ceramic in this experiment was less than that of the cured pure resin and hybrid polymer-ceramic material, which could explain why the cell viability of HGFs grows in direct contact with the ceramic was lower than that of the other groups. According to the ISO 10993-5:2009 standard, if the cell activity of the samples in the experimental group is more than 70% of that in the blank control group, it is considered to be non-cytotoxic. Therefore, the hybrid polymer-ceramic material has no cytotoxicity and meets the requirements for clinical use. The polymer has no cytotoxic effect on HGFs over a short time.
Nonetheless, the elution of polymer monomers into the oral cavity is a non-linear relationship with time [62]. Monomers can be released in trace amounts for a long time, affecting pulp viability and the prognosis of restored abutment. Furthermore, under the action of enzymes in saliva, monomeric degradation products regulate the growth of Streptococcus mutans and Streptococcus salivarius [63,64]. Therefore, it is necessary to further study the released elements and the content of monomers to evaluate the long-term safety.
In vivo experiments are necessary to demonstrate the effect of an actual clinical application.

Conclusions
Base on the hybrid polymer-ceramic dental material prepared in this experiment, the following conclusions can be drawn: (1) The sintering temperature and BAG content affect the mechanical properties of hybrid polymer-ceramic materials; (2) The mechanical properties of hybrid polymer-ceramic materials are similar to those of natural teeth; (3) A short-term exposure of human gingival fibroblasts to the hybrid polymer-ceramic material does not cause cytotoxicity. Institutional Review Board Statement: Not applicable.